This subject matter disclosed herein relates generally to gamma ray detectors, and more particularly, to systems and methods for gain calibration of gamma ray detectors.
Gamma ray detectors may be used in different applications, such as in Positron Emission Tomography (PET) systems. PET systems perform nuclear medicine imaging that generates a three-dimensional image or picture of functional processes within a body. For example, a PET system generates images that represent the distribution of positron-emitting nuclides within the body of a patient. When a positron interacts with an electron by annihilation, the entire mass of the positron-electron pair is converted into two 511 keV photons. The photons are emitted in opposite directions along a line of response. The annihilation photons are detected by detectors that are placed along the line of response on a detector ring. When these photons arrive and are detected at the detector elements at the same time, this is referred to as coincidence. An image is then generated based on the acquired image data that includes the annihilation photon detection information.
In PET systems, the gamma rays are detected by a scintillator in the scanning system, creating light that is detected by a photo-sensor (e.g., a photomultiplier tube (PMT), a silicon avalanche photodiode or a solid state photomultiplier). PET detectors based on vacuum photomultiplier photo-sensors require gain/energy calibration in order to properly operate. When using a small crystal array (e.g., 4×4 array) on solid state photomultiplier based PET detectors with multi-anodes (e.g., six anodes in a 2×3 array), it is often difficult to define reference crystals as the arrangement is very sensitive to the relative positioning of crystals over the anodes. Additionally, as the number of anodes increases, for example when six or more anodes are used, it is more difficult to apply conventional PET iterative algorithms to perform gain and energy calibration. For example, the processes may be more complex and time consuming.
For the case of a one-to-one coupling between the photo-sensor of solid state photomultiplier based PET detectors and the detector crystal (e.g., a Cerium doped Lutetium Yttrium Orthosilicate (LYSO) crystal), the 511 keV energy peak or known LYSO intrinsic background peaks can be used for calibration. However, in some designs, for example, a light-sharing block design among multiple anodes, this is not possible. Additionally, typical gain calibration for solid state photomultiplier based PET detectors may be performed by measuring a single photon pulse height when dark counts/current is small. To perform these measurements, a very high gain/low noise amplifier is needed. However, when the size of the photo-sensors in these solid state photomultiplier based PET detectors increases, for example, larger than 3×3 mm2 or when generating greater than 1 million dark counts per second, these measurements also cannot be performed because the single photon pulse cannot be identified properly due to count pileup. For example, in a 4×6 mm2 device, more than 10 million counts per second may be detected.
Thus, known processes for gain or energy calibration may not work satisfactorily for some configurations of gamma detectors, such as gamma detectors used in combination with photo-sensors, such as in solid state photomultiplier based PET detectors.